Flexible and Highly Biocompatible Nanofiber ... - ACS Publications

Flexible and Highly Biocompatible NanofiberBased Electrodes for Neural Surface Interfacing Dong Nyoung Heo,†,‡ Han-Jun Kim,§ Yi Jae Lee,⊥ Min Heo,‡ Sang Jin Lee,‡ Donghyun Lee,‡ Sun Hee Do,*,§ Soo Hyun Lee,*,⊥ and Il Keun Kwon*,‡ †

Department of Mechanical and Aerospace Engineering, The George Washington University, Washington, DC 20052, United States Department of Dental Materials, School of Dentistry, Kyung Hee University, Seoul 02447, Republic of Korea § Department of Clinical Pathology, College of Veterinary Medicine, Konkuk University, Seoul 05029, Republic of Korea ⊥ Center for BioMicroSystems, Korea Institute of Science and Technology, Seoul 02455, Republic of Korea ‡

S Supporting Information *

ABSTRACT: Polyimide (PI)-based electrodes have been widely used as flexible biosensors in implantable device applications for recording biological signals. However, the long-term quality of neural signals obtained from PI-based nerve electrodes tends to decrease due to nerve damage by neural tissue compression, mechanical mismatch, and insufficient fluid exchange between the neural tissue and electrodes. Here, we resolve these problems with a developed PI nanofiber (NF)-based nerve electrode for stable neural signal recording, which can be fabricated via electrospinning and inkjet printing. We demonstrate an NF-based nerve electrode that can be simply fabricated and easily applied due to its high permeability, flexibility, and biocompatibility. Furthermore, the electrode can record stable neural signals for extended periods of time, resulting in decreased mechanical mismatch, neural compression, and contact area. NF-based electrodes with highly flexible and body-fluid-permeable properties could enable future neural interfacing applications. KEYWORDS: nerve electrode, flexible device, neural interfacing, electrospun nanofiber, inkjet printing

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Although noninvasive nerve electrodes cause less discomfort than invasive electrodes after implantation, they also have a few drawbacks that must be mitigated. When the electrodes are placed onto neural tissue to precisely record neural electronic signals via the direct contact between the electrode and nerve, the relatively stiff structure of the electrodes can cause biological issues, such as inflammation, fibrous tissue formation, blood vessel compression, and neurological atrophy.13,14 To minimize these shortcomings, many studies have dedicated significant efforts in recent years to develop different electrode designs. Table 1 summarizes recently developed neural electrodes.15−26 There are two major avenues in the development of biocompatible electrodes: one is using different materials to produce softer substrates, and the other is revising the electrode design to minimize neural tissue compression and discomfort. With respect to mechanical properties, less stiff materials, such as

europrostheses linked with the nervous system have revolutionized strategies for treating patients with neurological disorders, such as spinal cord injury, brain stroke, and degenerative diseases, by forming electrical connections for recording and stimulating nervous system activity.1−4 In these systems, neural electrodes can be classified into two main groups, that is, invasive microneedle-type electrodes, which penetrate nerve fascicles, and noninvasive cuff-type electrodes, which are placed around the nerve trunk.5,6 Microneedle-type electrodes, such as the Michigan probe7,8 and the Utah array,9,10 have the advantages of recordings with increased signal-to-noise ratios (SNRs), enhanced signal selectivity, and reduced stimulus intensity. These electrodes also have some serious drawbacks, such as invasion-induced neural tissue damage after implantation. The initial damage is compounded by the presence of a deeply penetrating electrode. Therefore, cuff electrodes, which are noninvasive implantable neural prosthetics, are preferable for successful interfacing with the peripheral nervous system because they cause less tissue damage and allow longer neural communication.11,12 © 2017 American Chemical Society

Received: December 14, 2016 Accepted: February 14, 2017 Published: February 14, 2017 2961

DOI: 10.1021/acsnano.6b08390 ACS Nano 2017, 11, 2961−2971

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ACS Nano Table 1. Comparison of Recently Developed Neural Electrodes types of neural electrode for central nervous system

for peripheral nervous system

materials of substructure

shape

Young’s modulus

impedances at the biologically relevant frequency of 1 kHz

platinum−silicon composite carbon black

dura mater

10 MPa

5.2 kΩ15

fiber probe

2.38−3.0 GPa

1.3−2.8 MΩ16

thin film array helically wrapped ribbon

4 GPa 2.5 GPa

4.1−11.5 kΩ at 10 kHz17 6.2 kΩ18

C shape

8.45 GPa

11.1 kΩ19

polyimide (PI2611)

platinum gold nanoparticle and carbon nanotube gold and carbon nanotube titanium and gold

8.45 GPa

13.5 kΩ20

parylene C

platinum

4 GPa

8.3 kΩ21

PDMS

polypyrrole/polyol borate composite gold platinum graphene PEDOT:PSS and graphite−PDMS

helically wrapped sling shape self-locking with a strip and locking loop thin film array

68.9 MPa

∼10 kΩ22

microcable shape 3D sheath probe thin film array thin film array

1.81 MPa 4 GPa 4 GPa ∼1.0 MPa

100 kΩ23 20−60 kΩ24 243.5 kQ25 1.2 MΩ26

PDMS polycarbonate and cyclic olefin copolymer parylene C polyimide (durimide) polyimide (PI2611)

for brain nervous system

materials for sensing

PDMS parylene C parylene C PDMS

Figure 1. NF-based electrode with a patterned conductive layer. (a) Schematic illustration of the fabrication process using an inkjet printing system. (b) Concept of preproduction design and (c) image of the fabricated NF-based neural electrode. Electric wire cables were bonded to the end of the electrode pads using conductive silver epoxy and were covered with bone cement. (d−g) SEM images showing morphological differences among (d) Ag-patterned electrode, (e) PI NFs fused with Ag, (f) magnified image of PI NF with Ag, and (g) electrode after PEDOT growth. Scale bars for SEM images: (d) 1000 μm, (e) 50 μm, (f) 20 μm, and (g) 20 μm.

polyimide (PI), parylene C, and polydimethylsiloxane (PDMS), are usually used for the development of flexible electrodes instead of rigid silicon substrates. The Young’s moduli of silicon,27 PI,28 SU-8,29 parylene C,30 and PDMS23 are ∼25.5 GPa, 8.45 GPa, 5.6 GPa, 4.0 GPa, and 1.0 MPa, respectively. As a consequence,

compared with silicon-based electrodes, these electrodes induce less neural tissue damage and allow more stable neural communication. Additionally, these materials can be used to fabricate different designs with various shapes to reduce severe compression on the nerve tissue after implantation. Xiang et al. 2962


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Figure 2. Characteristic analysis of PI NFs as a printing substrate. (a) Surface morphology SEM images of PAA NFs (top) and PI NFs (bottom). (b) ATR-FTIR spectra of PAA and PI surfaces. (c) Strain−stress curves of the PI film and PI NFs. (d) Cross-sectional SEM images of PI NFs before (top) and after (bottom) applying pressure. (e) Incremental inclusion and (f) cumulative intrusion versus pore diameter of PI film and PI NFs before and after applying pressure, as measured using mercury. Scale bars for SEM images: (a) 50 μm, (b) 50 μm, (c) 500 μm, and (d) 200 μm.

platinum, titanium, and iridium) have a stiffening effect and prevent the neural electrode from fully conforming to the nerve trunk.35 Third, increasing the surface contact area between the implanted electrode and the neural tissue increases the nerve compression, resulting in nerve ischemia, Schwann cell necrosis, and nerve fiber degeneration.36 To overcome the drawbacks described above, we report the development of flexible and biocompatible nerve electrode consisting of an electrospun nanofiber (NF) membrane as the electrode substrate, silver nanoparticles (AgNPs) as the electronic connection pad, and poly(3,4ethylenedioxythiophene):poly(styrenesulfonate) (PEDOT:PSS) as an advanced conductive layer, and we demonstrate its capability of long-term in vivo neural signal recording and its stability for neural interfacing.

reported successful neural recording and stimulation from sciatic nerve branches using PI-based neural electrodes in the shape of helically wrapped ribbons.18 The authors demonstrated that the ribbon electrode showed the best performance and did not decrease blood flow in the connective vessels of the sciatic nerve after implantation. Xue et al. fabricated C-shaped cuff electrodes with gold carbon nanotubes on a PI substrate (∼12 μm) to improve the nerve−electrode interaction with minimum contact, thereby minimizing nerve tissue damage.19 Yu et al. reported a parylene-based, self-locking cuff electrode consisting of a strip and a locking loop that was easy to implant and did not induce nerve fascicle shrinkage over an 11 week implantation period.21 To obtain more flexible cuff electrodes and minimize the differences in mechanical strength between physical contact areas, Guo and McClain et al. developed a PDMS-based nerve electrode that reduced the difference in stiffness compared to neural tissue and showed electrophysiological recording functionality.22,23 However, even though the recently developed neural electrodes are technically flexible devices, several drawbacks still exist for a variety of reasons. First, neural electrodes made of flexible, polymer-based substrates (in the range of Epolymer substrate = 1.0 MPa to 8.45 GPa, including PDMS, parylene C, SU-8, and PI) are still rigid compared with soft neural tissue (Ebrain = 2.7− 3.1 kPa,31 Espinal cord = 3−6.3 kPa,32 Eperipheral nerve = 576−840 kPa33,34). Second, metal components that serve as the conductive layer embedded inside the polymer substrates (in the range of Emetal = 74−530 GPa, including gold, stainless steel,

RESULTS NF-Based Nerve Electrode Fabrication. In this study, we used an inkjet printing technique to pattern the recording sites of NF-based nerve electrodes due to its ability to simply and quickly manufacture samples. Through this technique, AgNPs could be patterned on the surface of PI NFs with a fine spatial resolution. After the AgNP patterns were printed on the surface of PI NFs or office paper, the resulting printing resolutions were 304.8 ± 17.1 and 362.1 ± 10.0 μm for the 10 μm input AutoCAD design, respectively, indicating better resolution on the PI NF surface than on office paper (Figure S1). In addition, the electric sheet resistance was observed to linearly decrease with the applied 2963


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Figure 3. Electrochemical properties and acute recording ex vivo. (a) Electrochemical impedance spectrum over a frequency range of 1−105 Hz; the impedance of the control electrode was 922.5 ± 57 Ω at 1 kHz, which decreased to 125.3 ± 1.4 Ω for the AgNP-coated PI NFs and to 123.7 ± 1.2 Ω for the AgNP/PEDOT-coated PI NFs. (b) Cyclic voltammograms with a scan range of −0.7 to +0.8 V at a scan rate of 100 mV/s; the charge delivery densities of the control electrode, AgNP-coated PI NFs, and AgNP/PEDOT-coated PI NFs were 0.17 ± 0.01, 89.53 ± 0.7, and 188.77 ± 3.54 mC/cm2, respectively (n = 3). (c) Image of a wrapped stimulus cuff electrode and a recording NF-based electrode. Acute ex vivo recordings obtained using (d) control electrode, (e) AgNP-coated PI NF electrode, and (f) AgNP/PEDOT-coated PI NF electrode.

AgNP printing number due to the increased accumulation of the conductive component. After a one-time patterning on the PI NFs, the heat-treated AgNP ink yielded a high sheet resistance; however, the sheet resistance of the AgNP-patterned PI NFs significantly decreased with increasing printing number, reaching as low as ∼0.31 Ω/sq for six printing repetitions (Figure S2). After the AgNP patterns were characterized, NF-based nerve electrodes were fabricated via inkjet printing technology. Figure 1a shows a schematic diagram of the fabrication process. First, PI NF sheets were placed on the paper feed tray and printed with AgNP ink in the designed patterns. The NF-based nerve electrodes consist of four main componentsstrips, locking holes, recording pads, and connection pads linked with electric wire (Figure 1b,c). Similar to tying a knot, locking structures with strips and locking holes allow the NF-based nerve electrode to be easily implanted into the peripheral nerve tissue, and its cuff diameter can be adjusted through the passed strip length. Based

on the preproduction design, AgNPs were patterned on the PI NF surface by inkjet printing (Figure 1d). After printing, the AgNP-patterned PI NFs were heat-treated at 140 °C to coalesce the AgNPs (Figure 1e,f), which forms electrical connections on the PI NFs. As shown in the scanning electron microscopy (SEM) images, printed AgNPs were melted and fused with random fiber structures by this heat treatment. Next, the conductive polymer PEDOT:PSS was deposited on the AgNPprinted PI NFs by electrochemical polymerization, which typically produces a thin layer with a highly uniform distribution and packing density (Figure 1g).37,38 Since the PEDOT:PSS was directly electropolymerized on printed AgNP, instead of nonconductive PI NFs, the growth method for the PEDOT:PSS was different than that described previous works, for example, conducting polymer nanotubes using poly(L-lactic acid) or poly(lactic-co-glycolic acid) NFs as PEDOT:PSS growth templates 39,40 and enzyme glucose oxidase-incorporated 2964


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Figure 4. In vivo neural signal recordings obtained using NF-based electrodes. (a) Electrical packaging electrode (left) and implanted electrode (right). The NF-based electrode is wrapped around the sciatic nerve of a rat. (b) Representative neural signal recording obtained from sciatic nerve tissue over a period of 12 weeks using a control electrode, an NF-based electrode, and an NF-based electrode soaked with tranilast. (c) Simultaneous mapping of the SNR of an NF-based electrode with and without tranilast.

40.3 ± 0.8% and a low tensile modulus of 41.0 ± 4.1 MPa. In addition, while there were no significant differences in the mechanical properties of the PI NFs before and after AgNP printing (Figure S3b), compared with the PI film, the control electrodes composed of the PI film substrate and Pt electrode contacts exhibited increased mechanical properties, due to the presence of a rigid metal (Figure S3a,c and Table S1). After thermal imidization, the nonwoven PI NFs were mechanically pressed by a heat transfer press machine to smooth the surface and reduce roughness, which provided high-quality printouts. As the pressure was applied, the thickness of the PI NFs was reduced from 434.2 ± 18.6 to 136.3 ± 10.8 μm, which can be observed in the cross-sectional SEM images in Figure 2d. Additionally, the intrusion rates of the PI film and PI NFs were studied to compare the changes in porosity according to the difference types of PI formation (Figure 2e,f). Compared with the unpressed PI NFs, the porosity of the pressed PI NFs decreased from 81.2 to 74.0%. Although the porosity of the PI NFs decreased after pressure was applied, the porosity remained higher than that of the PI film. Electrochemical Properties of the NF-Based Nerve Electrodes. The electrical properties of the fabricated NF-based nerve electrodes were evaluated by comparing its electrochemical impedance and charge delivery capacity (CDC) with those of a control electrode (i.e., a conventional cuff electrode), AgNP-coated PI NFs, and AgNP/PEDOT-coated PI NFs (n = 3). As shown in Figure 3a, the electrochemical impedance of AgNP/PEDOT-coated PI NFs was lower than that of the control electrode and the AgNP-coated PI NFs in the low-frequency range up to 2.3 kHz. Above this frequency, it exhibited similar values as compared with AgNP-coated PI NFs. The impedance of the control electrode, AgNP-, and AgNP/PEDOT-coated PI NFs at 1 kHz were 922.5 ± 57, 125.3 ± 1.4, and 123.7 ± 1.2 Ω, respectively. This difference was due to the enlarged surface activation area caused by PEDOT:PSS electropolymerization. In the cyclic voltammogram (CV), the electropolymerized PEDOT:PSS affected the CDC (Figure 3b). The calculated CDCs of the control electrode, the AgNP-coated PI NFs, and the AgNP/PEDOT-coated PI NFs were 0.17 ± 0.01, 89.53 ± 0.7,

PEDOT electrode formation.41 However, the accumulated PEDOT:PSS formed on the printed AgNP either covered the PI NF or grew around the PI NF. The use of PEDOT:PSS was necessary to ensure an enhanced electrical connection for successful nerve signal recording. Finally, electric wire cables were bonded to the end of the electrode pads using a conductive silver epoxy and were covered with bone cement. Characterization of the NF Membrane as an Electrode Substrate. Before the AgNPs were patterned via inkjet printing, we fabricated the PI NFs as a printing substrate and characterized their morphological, chemical, mechanical, and physical properties. PI NFs can be obtained from the precursor of PI, poly(amic acid) (PAA) NF by thermal imidization. The differences of their morphologies and chemical properties were characterized to confirm successful imidization. The PAA and PI NFs were randomly interconnected and nonwoven, with average diameters of 1023 ± 272 and 892 ± 264 nm, respectively (Figure 2a). The chemical differences between the PAA and PI NFs were confirmed using attenuated total reflectance Fourier transform infrared (ATR-FTIR) spectroscopy, as shown in Figure 2b. FTIR results of the PAA NFs showed a peak at 3430 cm −1 corresponding to the amine and carboxylic acid groups, a peak at 1648 cm−1 corresponding to amide I, and a peak at 1538 cm−1 corresponding to amide II. However, these peaks disappeared and new peaks appeared at 1776 cm−1 (CO symmetric stretching), 1371 cm−1 (C−N stretching), and 723 cm−1 (CO bending) in the spectrum of the PI NFs. Following the synthesis and characterization of PAA and PI, the PI NFs were compared and analyzed with PI film fabricated via solvent casting. To investigate the differences in the mechanical properties according to the type of PI formed, the PI film and PI NFs were subjected to strain−stress testing in uniaxial extension mode. As shown in Figure 2c, there were significant differences in the tensile properties of the PI film and the PI NFs due to their morphological differences. The PI film exhibited a high stiffness, with a low elongation at break of 12.1 ± 3.5% and a high tensile modulus of 1723.1 ± 325.3 MPa. However, the PI NFs had more flexible and elastic properties, with a high elongation at break of 2965


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Figure 5. Histological analysis of sciatic nerves after 2, 4, and 12 weeks of implantation. (a) Histological images of experimental groups. The structure of myelinated axons was better maintained in the NF-based electrode than in the PI film-based electrode group. In addition, the NFbased electrode with tranilast had less fibrous tissue infiltration than did the NF-based electrode. (b) Axonal indices (nerve area, axonal density, and total axonal number) of the experimental groups (black, control; red, NF-based electrode; green, NF-based electrode with tranilast). Throughout the experimental period, axonal densities were significantly greater in the NF-based electrode groups than in the PI film-based electrode group. Scale bars for histological images: 500 μm (low magnification, 4×), 50 μm (high magnification, 40×), and 20 μm (LFB-CEV, 100×); *p < 0.05 and **p < 0.01, compared with the control.

and 188.77 ± 3.54 mC/cm2, respectively. Meanwhile, Abidian et al. reported delamination of the electropolymerized PEDOT:PSS nanotube after five cycles of CV scanning.42 The AgNP/PEDOT-coated PI NF electrodes maintained their structure without delamination for up to 10 cycles of CV scanning despite slightly increased impedance due to AgNP oxidation (Figure S4). In addition, the functionalities of the NFbased nerve electrodes were tested on a sciatic nerve excised from a Sprague−Dawley (SD) rat. After cuffing the NF-based nerve electrode on the sciatic nerve tissue, it was wrapped such that the inner diameter was 1.7 mm (Figure 3c). All the recorded evoked nerve compound action potentials induced by an applied biphasic pulse current (peak potential = 300 μA, pulse duration = 100 μs) clearly showed the stimulus artifact signal followed by the evoked nerve signal. Although the evoked nerve compound action potential of the AgNP/PEDOT-coated PI NF electrode (0.768 mV, Figure 3f) to the applied electrical stimulation was relatively lower than that of the control (1.24 mV, Figure 3d), the evoked nerve compound action potential of the AgNP-coated PI NF electrode (0.151 mV, Figure 3e) was dramatically increased by the electropolymerized PEDOT:PSS. However, the evoked potential pulse duration between the artifact signal and the nerve signal of the NF-based electrodes was slightly delayed, which might be caused by gaps among the line electrodes due to the softness of the NF sheet. In Vivo Neural Signal Recording. Next, we implanted the NF-based nerve electrodes on the surface of the sciatic nerve of an SD rat and performed in vivo neural signal recordings for 12 weeks. The fabricated NF-based nerve electrode was connected

with an electric wire cable and covered with bone cement. The in vivo neural signal recording setup is shown in Figure 4a. During the implantation of the NF-based nerve electrode around the nerve, the lengths of the two strips were adjusted to the locking holes to wrap the electrode around the nerve circumference. For comparison, a PI film-based electrode (i.e., a conventional cuff electrode) was also prepared. In addition, the NF-based nerve electrode was soaked in tranilast solution as an antifibrotic agent43,44 to demonstrate the drug-loading ability and to enhance the long-term biocompatibility of the electrode with the sciatic nerve tissue. After the various electrodes were implanted, the nerve was stimulated through the soles of the rat’s foot using a mechanical stimulator. The ongoing electrical signal from the sciatic nerve is commonly generated by the activity of multiple electroneurograms (ENGs). As a consequence of mechanical stimulation, the continuous ENG signal can be recorded through the implanted electrode, and the recorded data were used without noise cancelation to compare the SNRs of the three electrodes. As shown in Figure 4b, all of the implanted electrodes could record differential ENG signals during the initial recording period. However, after 4 weeks of implantation, the control electrode (PI film-based electrode) showed no neural signal recording compared with the NF-based nerve electrodes, and this was due to the shrinkage of the nerve tissue and a newly formed fibrous tissue layer covering the electrode. This fibrous tissue layer was induced by the foreign body response and reduced the neural signal transfer from the tissue to the electrode (Figure S5). However, this biological issue can be reduced by decreasing the contact area between the tissue and the electrode. 2966


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Figure 6. Double immunofluorescence staining for MBP and S100B in the sciatic nerves at 2, 4, and 12 weeks postimplantation. Double immunofluorescence staining with antibodies against MBP (green) and S100B (red). The sections were counterstained with DAPI (blue). The numbers of S100B (Schwann cell) and MBP (myelin) double-positive axons were found to increase in tranilast-treated PI NF electrodes as compared to PI NF electrodes at 12 weeks. Scale bars = 25 μm.

When the surface area of the nerve electrode is decreased, the shrinkage of the nerve tissue resulting from neural compression decreases (Figure S6). According to the decreased contact area, the NF-based nerve electrodes were able to record stable ENG signals for 12 weeks. The SNRs measured from the NF-based nerve electrode with and without tranilast are summarized and shown as a heap map, which was generated from raw data recorded over a period of 12 weeks (Figure 4c). There were no significant differences among the other recordings, except for rat #6; this rat was implanted with an NF-based nerve electrode with tranilast and died from unknown causes. Additionally, the ENG signal qualities of both groups were almost equal throughout the measurement period. Therefore, we believe that the tranilast that was loaded on the NF-based nerve electrode to reduce fibrous tissue formation did not have any influence on the signal quality. To further evaluate the influence of tranilast, the histomorphological changes of the sciatic nerves were examined. As shown in Figure 5a, the degree of nerve degeneration differed among the experimental groups as time progressed. Axonal degeneration, characterized by edema, intraneural inflammatory cell infiltration, and axonal atrophy, was prominent in the PI film-based electrode groups at 4 weeks. In contrast, the density and total number of LFB-CEV-positive myelinated axons in the NF-based electrode groups (both with and without tranilast) were significantly higher than those in the PI film-based electrode group at 4 weeks (Figure 5b). Furthermore, the NF-based electrode groups with or without tranilast showed relatively constant nerve areas and axonal densities throughout the experimental period. In particular, for the tranilast-treated NFbased electrode, fibrous tissue infiltration was reduced, and the difference in the axonal indices was much more significant in the PI film-based electrode group than the NF-based electrode group up to 4 weeks. Consequently, the protein expressions which related to the axonal maintenance (Schwann cell and myelin) were analyzed using double immunofluorescence staining (Figure 6). Most of the axons showed staining for S100B (red) and MBP (green) in all experimental groups at 2 weeks. However, like the HE and LFB-CEV staining results, the number of S100B/MBP double-positive axons in the PI film-based

electrode decreased relative to those in the NF-based electrode groups (both with and without tranilast) at 4 weeks. Furthermore, the numbers of S100B and MBP double-positive axons in the tranilast-treated NF-based electrode were higher than those in the NF-based electrode at 12 weeks. These results indicate that the use of tranilast on NF-based electrodes affects biocompatibility, such that less fibrous tissue infiltration occurs and axonal structures are maintained. In addition, these findings may be indicative of the drug-loading capacity of the NF-based electrodes.

DISCUSSION The continuous monitoring of neural signals, which illustrates the function of neural tissue, is an important part of treating patients with traumatic neural disorders.12,45 To acquire stable neural signals for long periods of time, neural electrodes are required to have mechanical flexibility and conformability with respect to the soft nerve tissue, as these properties enable their long-term biocompatibility and reliability.35,46,47 According to these requirements, many researchers have developed highly flexible nerve electrodes to replace the conventional noninvasive nerve electrode composed of stiff substrates and rigid metal electrodes (Table 1). Although recently developed nerve electrodes are mechanically adaptive to soft nerve tissue, in contrast to conventional electrodes, these flexible electrodes still contain mechanical mismatches due to the presence of the conductive metal components that are embedded in the electrode substrate. In addition, they are still characterized by high electrochemical impedance values, which make them electrically active, although not sufficiently active to transduce high-quality neural signals. Moreover, no one has considered the permeability of the electrode substrate in the context of nutrient and oxygen exchange. As a long-term neural signal recording electrode, the substrate must have sufficient permeability to allow for the exchange of fluids, such as nutrients and oxygen, to support cell viability and thus to prevent the shrinkage of nerve tissue by promoting fluid retention.48,49 The ideal neural electrode should be able to exhibit the following main characteristics: (1) possess long-term, stable electrochemical 2967


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based nerve electrode is remarkably simple compared with the multistep fabrication process of conventional cuff electrodes. (ii) Printed conductive inks have better electrochemical properties than patterned metal thin films, and (iii) they penetrate the surface of porous NF membranes without significantly increasing the mechanical properties of the conductive metal ink-printed NFs. In addition, the NF-based nerve electrode provides (iv) decreased mechanical mismatch, (v) less neural tissue compression, (vi) increased porosity, and (vii) lower neural tissue contact area, thereby reducing nerve ischemia, Schwann cell necrosis, and nerve fiber degeneration. With these advantages, the NF-based nerve electrode is a promising electrode for use in long-term bio-interconnections and the continuous monitoring of neural signals linked with the peripheral nervous system. Although additional technological developments are necessary to electrically stimulate and recode selective nerve fibers, these findings will be helpful for establishing highly flexible, body fluid-permeable, stable nerve electrodes for neural interfacing applications.

properties; (2) induce minimal nerve atrophy; and (3) allow continuous neural signal recording. To develop a flexible nerve electrode that meets these requirements, an NF-based nerve electrode was designed using both electrospinning and inkjet printing technologies. The results of this study demonstrate that the designed NFbased nerve electrode has electrochemical properties superior to those of conventional cuff electrodes (Figure 3) and that it is able to record neural signals for a long period of time following implantation (Figure 4). In addition, our NF-based electrode design provides (i) an adjustable cuff that can be fitted to wrap around the nerve circumference, (ii) close contact between the recording sites and the circular nerve trunk, and (iii) optimal spatial distribution for stable, long-term neural signal recording. After implantation, the conventional cuff electrode caused immune-mediated pathological tissue reactions, such as inflammation and neural degeneration, due to its mechanical mismatch with nerve tissue, its nonpermeable structure, and neural compression. Unlike conventional cuff electrodes, the NFbased nerve electrode enables the delivery of nutrients and oxygen to the peripheral nerve tissue due to its porous structure. Furthermore, the many open pores of the NF-based nerve electrode decreased the mechanical mismatch with the nerve tissue as well as the contact area with the surface of the peripheral nerve after implantation (Figure 2 and Figure S6). As a consequence, the nerve trunk was under less neural compression without loose contacts between the electronic connection pads and the nerve tissue, resulting in increased long-term biocompatibility and decreased nerve atrophy (Figures 5 and 6). In addition, the NF-based nerve electrode could be loaded with tranilast to promote antifibrotic and anti-inflammatory effects. The tranilast-loaded, NF-based nerve electrode was able to decrease neural degeneration and support a healthy axonal density. In paper-based electronics, metal inks composed of conductive materials are usually patterned via inkjet printing.50,51 AgNPs have been widely used in such electronic devices due to their high conductivity, short manufacturing time, and cost-effective fabrication.52−55 However, to the best of our knowledge, the integration of conductive AgNPs and electrospun NF membranes via inkjet printing for the development of electronic devices has not yet been reported. Although these two main components (i.e., printable AgNP ink and flexible NF membranes) have been developed in their respective fields, certain factors inhibit their integration. The principal reason is that printed AgNPs should be heat-treated to melt and coalesce, forming Ag traces with good electrical connections; however, most NF membranes have low thermal resistances.56−58 Thus, if a heat treatment is applied to melt AgNPs on a conventional polymer-based NF membrane, the membrane would be burned out. Therefore, it is necessary to use a thermally resistant polymer for patterning conductive AgNPs. In this study, PI NFs were used as a substrate to develop the flexible electrode due to their excellent thermal, electrical, and mechanical properties as well as long-term stability.28,59,60 After inkjet printing, uniformly distributed AgNPs were successfully patterned on the PI NF surfaces without increasing any mechanical properties (Figure S3). Then, via the electric deposition of crystallized PEDOT:PSS as an electrical conducting component, an enhanced conductive layer was effectively formed inside the PI NF sheets (Figure 1). As described above, this form of NF-based nerve electrode has many advantages over conventional planar substrates patterned with conductive materials. (i) The fabrication process of the NF-

CONCLUSIONS In summary, we designed and prepared an innovative NF-based nerve electrode by combining electrospinning and inkjet printing technology. The NF-based nerve electrode can be simply fabricated and easily applied due to its high permeability, flexibility, and biocompatibility. Electrochemical analysis showed that the NF-based nerve electrode has more enhanced electrochemical properties than the conventional cuff electrode. After implantation, the NF-based nerve electrode is able to record neural signals for a long period of time. In addition, the NF-based nerve electrode showed reduced immune-mediated pathological tissue reactions, decreased nerve atrophy, and increased long-term biocompatibility. These results might be useful to improve long-term bio-interconnections and the continuous monitoring of neural signals linked with the nervous system. Furthermore, to our knowledge, this study is the pioneering application of an electrospun nanofiber that confirms its effect for recording biological signal recording in vivo. Our findings suggest that this can be widely applicable to the neural tissue engineering field. EXPERIMENTAL SECTION NF-Based Electrode Fabrication. Electrospun PI was fabricated by a three-step method. First, the precursor, PAA, was synthesized by the polycondensation of pyromellitic dianhydride (4.3624 g) and 4,4′oxidianiline (4.0048 g) in N,N-dimethylacetamide (30.6 g). Then, the PAA solution was loaded into a glass syringe and electrospun onto a rolling mandrel with high-voltage power (18 kV). Finally, the electrospun PAA NFs were thermally imidized at 350 °C for 1 h in an oven and subsequently cooled to room temperature, yielding PI NFs. To enhance the quality of AgNP patterning, the PI NFs were mechanically pressed on a heat transfer press machine at a high temperature of 120 °C for 30 min. After PI NF preparation, the AgNPs were patterned onto the substrate using an EPSON C88+ desktop printer containing a AgNP ink cartridge (JS-B25P, NovaCentrix Co.). The nerve electrode design for the AgNP patterning was easily input and printed using AutoCAD. The AgNP-patterned PI NFs were then heat-treated at 180 °C in air for 1 h to form highly conductive silver traces. Electropolymerization of PEDOT:PSS. PEDOT:PSS was polymerized electrochemically over four AgNP printed line electrode sites. The PEDOT:PSS solution was composed of 0.01 M 3,4-ethylenedioxythiophene (EDOT) (Sigma-Aldrich, St. Louis, MO, USA), 0.1 M poly(sodium 4-styrenesulfonate) (Sigma-Aldrich, St. Louis, MO, USA), and 100 mL of triple-distilled water. Electrochemical polymerization was performed by the galvanostatic mode using an Autolab 2968


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electrode, and tranilast-loaded, NF-based electrodes) were prepared for the in vivo study. Tranilast was loaded via a soaking and drying method.62 In brief, tranilast (2 mg) was dissolved in ethanol (1 mL), and the solution (200 μL) was spread onto the NF-based electrode and then air-dried. All electrodes were sterilized under a UV-sterilizing, laminar flow hood and exposed for 2 h to a UV-C germicidal lamp (UV output: 19.8 W, Sankyo Denki, Japan). The left sciatic nerve was exposed by an incision of the lateral surface of the hindlimb parallel to the femur, and the electrodes were implanted at the center of the sciatic nerve. All leads were routed subcutaneously to a pin connector for continuous data acquisition, which was sealed with dental cement (Vertex Self-Curing). Then, the functionality of the implanted cuff electrode was examined weekly by recording the afferent ENG signal of the sciatic nerve upon mechanical stimulation using a von Frey filament, representing a force of linear pressure corresponding to a mass of 75 g. Histological Analysis. Rats were sacrificed at 2, 4, and 12 weeks after implantation using CO2. Both the implanted cuff electrodes and the surrounding tissue were collected and fixed in 10% neutral-buffered formalin (BBC Biochemical, WA, USA). Then, the sciatic nerve was processed according to routine histological methods (n = 3 per each group and time point). The tissues were embedded in paraffin and cut in 4 μm thick sections for hematoxylin and eosin (BBC Biochemical) and Luxol fast blue-cresyl etch violet staining (LFB-CEV, American MasterTech, CA, USA). Quantitative data of the nerve area and axonal indices (axonal density and total axon number) were measured using Leica Application Suite software (Leica Microsystems, Wetzlar, Germany). Axonal densities were calculated by dividing the average number of LFB-CEV-positive myelinated axons in at least five sections by the area of the micrograph. The total number of myelinated axons in the tissue was quantified by multiplying the estimated axonal density by the nerve area. Double Immunofluorescence. Serial formalin-fixed tissue sections were used for double immunofluorescence staining. Briefly, the sections were deparaffinized and underwent heat-induced antigen retrieval using a combination of citric acid (Sigma-Aldrich, St. Louis, MO, USA) and trisodium citrate (Sigma-Aldrich) to liberate the crosslinked epitopes. The sections were then incubated in blocking solution (3% bovine serum albumin, Sigma) for 30 min. The primary antibodies used were polyclonal rabbit anti-MBP antibody (Wuhan Fine Biological Technology, Hubei, China, 1:100) and monoclonal mouse anti-S100B (Abcam, 1:100). After being washed with 1× PBS for 3 × 5 min, they were incubated with a combination of FITC-conjugated goat anti-rabbit and APC-conjugated goat anti-mouse secondary antibodies (Santa Cruz Biotechnology, 1:100) for 30 min. The sections were fixed with fluoromount using DAPI (Vector Laboratories, CA, USA) and analyzed under a fluorescence microscope (Leica Microsystems). Statistical Analysis. All values are expressed as mean ± standard deviations. Statistical analysis was performed using PASW Statistics 21 software (SPSS, Inc., Chicago, IL). Multiple comparisons were analyzed using one-way analysis of variance (ANOVA) followed by Dunnett’s T3 post-hoc paired comparison test. A value of p < 0.05 are considered to be statistically significant.

PGSTAT 302N (EcoChemie, Utrecht, The Netherlands). A conventional three-electrode configuration was used for the electrochemical polymerization of PEDOT:PSS in the galvanostatic mode. The working electrode was connected to the line electrode sites of the functional nerve cuff electrode through an external connector. A Ag/AgCl electrode was used as a reference electrode, and Pt wire was used as a counter electrode. A current density of 8 μA/mm2 was applied to each line electrode site for 300 s. Analytical Equipment. To examine the sample morphologies, all of the samples were sputter-coated with platinum for 10 min, and their surface and cross-sectional morphologies were imaged by SEM (Hitachi S-4700). The chemical structures of PAA and PI were characterized using an FTIR spectrophotometer (Spectrum One System, PerkinElmer). To estimate the mechanical properties, the specimens were prepared with dimensions of 20 mm × 5 mm (length × width). Their strain−stress curves were measured using a tensile testing machine (EZSX, Shimadzu) according to the ASTM D882 method. Porosity was measured using a porosimetry analyzer (AutoPore IV 9500, Micromeritics Co.) by applying various levels of pressure to a sample immersed in mercury. Electrochemical Characterization. The fabricated electrodes were investigated to determine their interfacial impedance and charge delivery capacity in a 0.1 M phosphate-buffered saline solution (pH 7.0) using electrochemical impedance spectroscopy and cyclic voltammetry. The electrochemical experiments of the electrodes were performed at room temperature using an Autolab system (PGSTAT 302N, NOVA software, Ecochemie, Utrecht, The Netherlands). A conventional threeelectrode system was used to determine impedance. The system included a fabricated electrode as a working electrode, a Pt wire as a counter electrode (0.5 mm in diameter, 30 mm in length, and 99.95% purity), and a Ag/AgCl reference electrode. An AC sinusoidal signal with an amplitude of 10 mV was applied to measure the impedance between the electrode and the electrolyte over a frequency range of 1− 105 Hz. The CV results were obtained after 20 repetitive potential scans to ensure that the double-layer capacitance of the electrode had stabilized. A scan rate of 100 mV/s was used, and the potential between the working electrode and the reference electrode was kept in the range of −0.65 to +0.8 V to avoid electrolysis of water. The CDC was calculated using the following equation:61

CDC =

1 vA

∫E

Ea

|i|dE (C/cm 2)

c

where v is the scan rate (V/s), A is the geometrical surface area of the electrode (cm2), Ea/Ec is the anodic/cathodic potential limit, i is the measured current, and E is the electrode potential (V vs Ag/AgCl). Acute Ex Vivo and In Vivo Testing. Ten week old male SD rats were housed and handled in accordance with the regulations of the Institutional Animal Care and Use Committee of Konkuk University (KU16049). For the acute ex vivo test, animals were sacrificed using CO2, and the sciatic nerves were collected. The dissected sciatic nerves were immediately placed in Krebs solution. A stimulating electrode (PI film-based) and a recording electrode (PI NF-based) were wrapped around the nerve with a 5 mm gap between the electrodes. The ground electrode was placed in a separate Krebs solution. The stimulating cuff electrode was connected to a pulse stimulator (isolated pulse stimulator, model 2100, A-M Systems, Sequim, WA, USA). Additionally, the recording cuff electrode was connected to a differential amplifier (differential AC amplifier, model 1700, A-M systems, Sequim, WA, USA). The amplified neural signals of the sciatic nerve were collected using a data acquisition device (NI USB-6356, National Instruments, Seoul, Korea). Then, the collected signals were processed using LabVIEW software (National Instruments, Seoul, Korea) and displayed on a laptop computer. The electrical stimulation was applied with a pulse amplitude of 300 μA and a pulse width of 100 μs. The differential amplifier gain was 1000. For the in vivo study, surgical procedures were performed under general anesthesia induced by xylazine HCl (Bayer Korea, Gyeonggi-do, Korea), and anesthesia was maintained with 1.5−2% isoflurane (with 100% O2). Three types of electrodes (control−PI film, NF-based

ASSOCIATED CONTENT S Supporting Information *

The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acsnano.6b08390. Additional information and data of optical images of AgNP patterns on office paper and PI NF surfaces and their diameters, sheet resistance of a AgNP-printed PI NF, strain−stress curves of the control electrode, PI film, NFbased electrode and PI NF, SEM images of AgNP/ PEDOT-coated PI NFs before and after 10 cycles of CV scanning, histological images of small intestinal submucosa wrapping around the sciatic nerve and ex vivo ENG signal recording, histological images of sciatic nerves after implantation of PI film-based electrodes with small and large holes (PDF) 2969


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ACS Nano

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AUTHOR INFORMATION Corresponding Authors

*E-mail: [email protected]. *E-mail: [email protected]. *E-mail: [email protected]. ORCID

Dong Nyoung Heo: 0000-0002-7717-7184 Author Contributions

D.N.H., H.-J.K, Y.J.L., M.H., S.J.L., D.L., and S.H.D. performed the research. D.N.H., H.-J.K., and Y.J.L. wrote the manuscript. D.N.H., S.H.D., S.H.L., and I.K.K. designed the research. S.H.D., S.H.L., and I.K.K. supervised the project. All the authors read and commented on the manuscript. Notes

The authors declare no competing financial interest.

ACKNOWLEDGMENTS This research was supported by the Public Welfare & Safety research program through the National Research Foundation of Korea (NRF) funded by the Ministry of Education, Science and Technology (NRF-2012R1A5A2051388), the Korea Health Technology R&D Project through the Korea Health Industry Development Institute (KHIDI) funded by the Ministry of Health & Welfare (HI14C2241), and the Basic Science Research Program through the NRF funded by the Ministry of Science, ICT & Future Planning (NRF-2014R1A1A3052557). This research was also partially supported by the KIST Institutional Program (2E26180). REFERENCES (1) del Valle, J.; Navarro, X. Interfaces with the Peripheral Nerve for the Control of Neuroprostheses. Int. Rev. Neurobiol. 2013, 109, 63−83. (2) Grill, W. M.; Norman, S. E.; Bellamkonda, R. V. Implanted Neural Interfaces: Biochallenges and Engineered Solutions. Annu. Rev. Biomed. Eng. 2009, 11, 1−24. (3) Rossini, P.; Burke, D.; Chen, R.; Cohen, L.; Daskalakis, Z.; Di Iorio, R.; Di Lazzaro, V.; Ferreri, F.; Fitzgerald, P.; George, M.; et al. NonInvasive Electrical and Magnetic Stimulation of the Brain, Spinal Cord, Roots and Peripheral Nerves: Basic Principles and Procedures for Routine Clinical and Research Application. An Updated Report from an I.F.C.N. Committee. Clin. Neurophysiol. 2015, 126, 1071−1107. (4) Micera, S.; Navarro, X. Bidirectional Interfaces with the Peripheral Nervous System. Int. Rev. Neurobiol. 2009, 86, 23−38. (5) Ortiz-Catalan, M.; Brånemark, R.; Håkansson, B.; Delbeke, J. On the Viability of Implantable Electrodes for the Natural Control of Artificial Limbs: Review and Discussion. Biomed. Eng. Online 2012, 11, 33. (6) Badia, J.; Boretius, T.; Andreu, D.; Azevedo-Coste, C.; Stieglitz, T.; Navarro, X. Comparative Analysis of Transverse Intrafascicular Multichannel, Longitudinal Intrafascicular and Multipolar Cuff Electrodes for the Selective Stimulation of Nerve Fascicles. J. Neural Eng. 2011, 8, 036023. (7) Fattahi, P.; Yang, G.; Kim, G.; Abidian, M. R. A Review of Organic and Inorganic Biomaterials for Neural Interfaces. Adv. Mater. 2014, 26, 1846−1885. (8) Cheung, K. C. Implantable Microscale Neural Interfaces. Biomed. Microdevices 2007, 9, 923−938. (9) Wark, H.; Mathews, K.; Normann, R.; Fernandez, E. Behavioral and Cellular Consequences of High-Electrode Count Utah Arrays Chronically Implanted in Rat Sciatic Nerve. J. Neural Eng. 2014, 11, 046027. (10) Christensen, M.; Pearce, S.; Ledbetter, N.; Warren, D.; Clark, G. A.; Tresco, P. A. The Foreign Body Response to the Utah Slant Electrode Array in the Cat Sciatic Nerve. Acta Biomater. 2014, 10, 4650− 4660. 2970


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